The present invention relates to improvements in and relating to magnetic resonance imaging (MRI) apparatus and methods, and in particular to a method and apparatus for the thermographic imaging of a subject, generally although not essentially a human or animal body, and to contrast agents and media for use in such methods.
In certain cancer treatments, malignant tissue within the body is destroyed by irradiation with radiation, for example microwave radiation, which has sufficient heating effect to kill the malignant tissue, e.g. by raising the local temperature to about 43.degree. C. As will readily be appreciated, the heating radiation can also kill healthy tissue and it is therefore of great importance for such treatments for the physician to be able to determine the temperature at and near the irradiated site.
This is particularly important since the radiation reflection and absorption characteristics are not uniform throughout the body and, especially where two or more directed radiation sources are used to achieve the heating effect, there is a danger that radiation reflection or shadowing by body tissue may cause areas of significant temperature increase ("hot-spots") to occur in healthy tissue or may prevent the temperature increase in part or all of the malignant tissue site from being sufficient to kill off all the malignant cells.
Several methods of temperature monitoring have been proposed, but to date all such methods have been either invasive, insufficiently accurate or time consuming or have enabled temperatures to be measured for superficial tissue layers only. Thus typical techniques which have been used include invasive monitoring by insertion of thermal sensing probes, infrared thermography, CAT scanning and NMR relaxation rate assessments.
There remains a need for a non-invasive thermographic imaging method capable of determining local temperatures throughout the body with reasonable accuracy.
We have now found that using a modification of our recently developed Electron Spin Resonance Enhanced Magnetic Resonance Imaging (ESREMRI) method thermographic imaging, or temperature monitoring, can be effected.
MRI is a diagnostic technique that has become particularly attractive to physicians as it does not involve exposing the patient to the harmful X-or gamma-radiations of conventional radiographic imaging techniques.
In our co-pending European Patent Application EP-A-296833 and British Patent Applications Nos. 8817137 and 8819753.8 we have described how the intensity of the magnetic resonance (MR) signal from which MR images are built up may be enhanced, e.g. by factors of 100 or more, by exciting an esr transition of a paramagnetic substance present within the subject being imaged where that esr transition is coupled to the nmr transition of the nuclei (generally protons and usually protons in water molecules) which emit the MR signals from which the MR images are built up.
The degeneracy of the spin states of nuclei with non-zero spin, e.g. .sup.1 H, .sup.13 C, .sup.19 F, etc., is lost when such nuclei are within a magnetic field and transitions between the ground and excited spin states can be excited by the application of radiation of the frequency (.omega..sub.o) corresponding to energy difference E of the transition (i.e. .omega..sub.o = E). This frequency is termed the Larmor frequency and is proportional to the strength of the applied field. As there is an energy difference between the spin states, when the spin system is at equilibrium the population distribution between ground and excited spin states is a Boltzmann distribution and there is a relative overpopulation of the ground state resulting in the spin system as a whole possessing a net magnetic moment in the field direction. This is referred to as a longitudinal magnetization. At equilibrium the components of the magnetic moments of the individual non-zero spin nuclei in the plane perpendicular to the field direction are randomized and the spin system as a whole has no net magnetic moment in this plane, i.e. it has no tranverse magnetization.
If the spin system is then exposed to a relatively low intensity oscillating magnetic field perpendicular to the main field and produced by radiation at the Larmor frequency, generally radiofrequency (RF) radiation in conventional MRI, transitions between ground and excited spin states occur. If the exposure is for a relatively short duration then the resultant magnitudes of the longitudinal and transverse magnetizations of the spin system are functions of the exposure duration which oscillate about zero at the Larmor frequency and are 90.degree. out of phase with each other. Thus, from equilibrium, a pulse of duration (2n+1).pi./2.omega..sub.o (a so-called 90.degree. pulse when n is even and a 270.degree. pulse when n is odd) leaves the system with maximum transverse magnetization (of magnitude proportional to the initial longitudinal magnetization at equilibrium) and no longitudinal magnetization, a pulse of duration (2n+1).pi./.omega..sub.o (a 180.degree. pulse) leaves the system with inverted longitudinal magnetization and inverted transverse magnetization (and hence from equilibrium no transverse magnetization), etc.
When the pulse is terminated, the oscillating magnetic field produced by any resulting net transverse magnetization can induce an oscillating electrical signal (of angular frequency .omega..sub.o) in a detector coil having its axis arranged perpendicular to the main field direction. For this purpose the transmitter used to emit the pulse can also be used as a detector.
Induced nuclear magnetic resonance signals, hereinafter termed free induction decay (FID) signals, have an amplitude proportional to the transverse magnetization (and hence generally to the original population difference between ground and excited spin states).
If the nuclei of the spin system experienced an entirely uniform magnetic field, the FID signal would decay due to spin-spin interactions at a rate with a characteristic time of T.sub.2, the transverse or spin-spin relaxation time. However, due to local field inhomogeneities, the nuclei within the spin system will have a spread of Larmor frequencies and decay of transverse magnetization is more rapid, having a characteristic time of T.sub.2 * where 1/T.sub.2 * =1/T.sub.2 +1/T.sub.inh, T.sub.inh representing the contribution due to field inhomogeneities. T.sub.2 itself can be determined using spin-echo imaging in which, after the decay of the FID signal (usually following a 90.degree. pulse) the system is exposed to a 180.degree. pulse and an "echo" signal is generated, the decay in the amplitude of the echo being governed primarily by T.sub.2 as, with the inversion of the transverse magnetization for the individual nuclei, the field inhomogeneities referred to above cause tranverse magnetization to build up to a maximum at time TE/2 after the 180.degree. pulse where the time between the previous maximum transverse magnetization and the 180.degree. pulse is also TE/2.
To generate different images, different pulse and FID detection sequences are used. Perhaps the simplest is saturation recovery (SR) where the FID signal is determined after a single 90.degree. initiating pulse. The signal strength is dependent upon the magnitude of the longitudinal magnetization before the pulse, and hence on the nuclear density and the extent to which the system reequilibrates in the time (TR) between successive initiating pulses. In spin-echo imaging, for example multiple-echo imaging, the pulse and detection sequence may be: initiating 90.degree. pulse (at time 0), FID detection (following the initiating pulse), 180.degree. pulse (at time TE/2), detection of 1st echo (at time TE), 180.degree. pulse (at time 3TE/2), detection of 2nd echo (at time 2TE) . . . , initiating pulse for the next sequence (at time TR), etc. In this technique, a TR is selected which is sufficient for a reasonable reequilibration to occur in the period between successive initiating pulses.
As is explained further below in connection with the example of two dimensional Fourier transformation (2DFT) image generation, in order to generate a single image with adequate spatial resolution, it is necessary to perform a large number (e.g. 64-1024) of separate pulse and detection sequences.
Since TR has in principle to be large with respect to T.sub.1, the characteristic time for relaxation of the excited system towards the equilibrium Boltzmann distribution between ground and excited spin states, to permit longitudinal magnetization to build up between successive pulse sequences so as to avoid the FID signal strength decaying in successive pulse sequences, the total image acquistion time is generally relatively large. Thus, for example, TR may conventionally be of the order of seconds and the image acquisition time may be of the order of 10-30 minutes.
Certain so-called fast imaging (FI) techniques may be used to accelerate reequilibration and so reduce image acquisition time; however they inherently result in a reduction in the S/N ratio and/or contrast hence in poorer image quality. The FI technique involves for example exciting the spin system with a less than 90.degree. pulse and thus the difference between ground and excited spin state populations is only reduced rather than eliminated (as with a 90.degree. pulse) or inverted and so reattainment of equilibrium is more rapid. Nevertheless, the transverse magnetization generated by the less than 90.degree. pulse is less than that for a 90.degree. pulse and so FID signal strength and thus S/N ratio and the spatial resolution in the final image are reduced.
Using different pulse and detection sequences and by manipulation of the acquired data, MRI can be used to generate a variety of different images, for example saturation recovery (SR), inversion recovery (IR), spin echo (SE), nuclear (usually proton) density, longitudinal relaxation time (T.sub.1) and transverse relaxation time (T.sub.2) images. Tissues or tissue abnormalities that have poor contrast in one such image often have improved contrast in another. Alternatively, imaging parameters (nuclear density, T.sub.1 and T.sub.2) for tissues of interest may be altered by administration of a contrast agent. Thus many proposals have been made for the administration of magnetically responsive materials to patients under study (see for example EP-A-71564 (Schering), U.S. Pat. No. 4,615,879 (Runge), WO-A-85/02772 (Schroder) and WO-A-85/04330 (Jacobsen)). Where such materials, generally referred to as MRI contrast agents, are paramagnetic (for example gadolinium oxalate as suggested by Runge) they produce a significant reduction in the T.sub.1 of the water protons in the zones into which they are administered or at which they congregate, and where the materials are ferromagnetic or superparamagnetic (e.g. as suggested by Schroder and Jacobsen) they produce a significant reduction in the T.sub.2 of the water protons, in either case resulting in enhanced (positive or negative) contrast in the magnetic resonance (MR) images of such zones.
The contrast enhancement achievable by such agents is limited by a number of factors. Thus such contrast agents cannot move the MRI signal intensity (I.sub.s) for any tissue beyond the maximum (I.sub.l) and minimum (I.sub.o) intensities achievable for that tissue using the same imaging technique (e.g. IR, SR, SE, etc.) in the absence of the contrast agent: thus if "contrast effect" is defined as (I.sub.s -I.sub.o)/(I.sub.l -I.sub.o), contrast agents can serve to alter the "contrast effect" of a tissue within the range of 0-1. However to achieve contrast improvement an adequate quantity of the contrast agent must be administered to the subject, either directly to the body site of interest or in such a way that the natural operation of the body will bring the contrast agent to that body site.
ESREMRI utilises the spin transition coupling phenomenon known in conventional nmr spectroscopy as the Overhauser effect to amplify the population difference between ground and excited nuclear spin states, producing a significant overpopulation (relative to the Boltzmann distribution population) of the excited spin state of the nuclear spin system producing the MR image. This is achieved by exciting a coupled esr transition in a paramagnetic species naturally occurring in or introduced into the sample being imaged, which is generally but not essentially a human or animal subject.
The MRI apparatus for use according to this technique requires a second radiation source for generating the radiation capable of stimulating such an esr transition as well as the first radiation source for generating the radiation used to stimulate the nuclear spin transition.